JP4971981B2 - Cell-independent production of tissue equivalents - Google Patents

Abstract

This invention relates to cell-independent processes which mimic cellular bioremodelling and produce organized biomaterials which have mechanical properties and viable cell densities suitable for use as functional tissue implants. The biomaterials are produced by providing a gel comprising a matrix of scaffold fibers of and an interstitial fluid; and plastically compacting the gel to produce the biomaterial. The biomaterials may comprise 3D structures such as layering, alignment and meso-scale zonal heterogeneities of cells and matrix which mimic native tissue structure. Biomaterials with biomimetic structure as described herein may be useful in a range of therapeutic applications.

Description

  The present invention relates to a method for producing a tissue equivalent implant for repairing and / or replacing damaged tissue in an individual, and to an implant produced by the method.

  Conventionally, tissue engineering aims to convert a scaffold structure of progenitor cells into a tissue-like structure having a biomimetic function. Such conversion processes generally involve cell-based remodeling in culture. However, in most cases, cell-based remodeling is time consuming (often weeks), difficult to control and expensive, and has a very limited ability to form bioartificial materials or “tissues”. (M. Eastwood et al., Cel. Motil. Cytoskel. 1998 40 13; D. Huang et al., Ann. Biomed. Eng. 1993 21 289). This is due, in part, to perfusion / hypoxia limitations related to tissue density (eg, ligament, skin and muscle tissue density). This situation is exacerbated by the poor understanding of how cells actually create specific natural microstructures (ie, three-dimensional cell matrix tissues). The material composition of bioartificial constructs, and more particularly the three-dimensional nano-micro (meso) scale structure, is crucial for its success (RA Brown, in Future Strategies for Tissue and Organ Replacement (JM Polak, LL Hench, P. Kemp), World Scientific Publishing, Singapore (2002) 48; RA Brown et al., Wound Rep. Reg. (1997) 5 212).

  Natural matrix scaffolds (eg, collagen, fibrin, hyaluronan, fibronectin) have been used and made to have a suitable biomimetic three-dimensional meso-scale structure for cell / tissue growth. However, progress has been limited due to the limited ability to control the construction (assembly) of biomimetic structures on this scale, especially with natural proteins and living cells. This is due in part to the limited knowledge of three-dimensional assembly of natural protein macromolecules (F. Volrath, DP Knight, Nature 2001, 410, 541) and the required assembly / remodeling process. It is necessary to rely on cells to carry out.

  Cell-seeded collagen gels are a good starting point for biomimetic but are very weak mechanically (Z. Feng et al., Artificial Organs 2003, 27, 84; L. Krishnan et al., Tissue Eng. 2004 10 241 ), It takes time to become stronger in culture.

  Conventional, cell-based fabrication examples utilize several systems. Most simply, such production is the development of a superconfluent two-dimensional culture that does not use any supporting scaffold at all (S. Calve et al., Tissue. Eng. 2004, 10, 755). Huang et al. (Huang et al., 1993, supra) used a collagen gel under uniaxial tension to create a ligament-like tissue that reached a fracture stress of 0.14 MPa after 12 weeks. Feng et al. (Feng et al., 2003 supra) found that unloaded and non-oriented collagen gels decreased in diameter by 85% during culture over 10 weeks and yield stress was approximately 0.1 MPa. Garvin et al. (Garvin, J. et al., Tissue Eng. 2003, 9, 967) reported using loads to provide cell / fibril alignment and an epitenon-like surface. The fracture stress of the 7-day construction increased from 0.11 to 0.33 MPa (modulus 1.8 MPa) at 1 h / day, 1% strain, 1 Hz cyclic loading.

  Insoluble collagen sponges (I. V. Yannas et al., Science (1982) 215, 174) (eg, Integra ™) have higher strength than collagen gels. Collagen sponges are random porous structures derived from extensively prepolymerized fibers, sheets or meshes, and are often derived from reconstituted or fragmented insoluble polymers (eg, skin, tendons, etc.) Animal collagen) from high temperature (ie well above the protein denaturation temperature; eg> 60 ° C.). Thus, the microstructure of collagen sponges is not biomimetic and is not easily remodeled by cells. A technique for strengthening insoluble collagen sponges by hot compression has recently been reported for drug delivery (Z. Rusczcak, W. Friess, Adv. Drug Delivery Rev. 2003 55 1679; US 2003095997; US 2003133967).

  Another practical alternative is the small intestinal submucosa (SIS) collagen membrane (S.F. Badylak et al., J. Biomed. Mater. Res. 1995 29 977). Although SIS has excellent mechanical and cell growth properties, it remains an animal product that has been harvested and processed. As such, it is cell-free and its collagen structure and three-dimensional tissue are predetermined by animal tissue.

  The present invention relates to the following surprising findings, in that physical processes that require little or no cellular involvement (ie, cell-independent processes) are effectively The finding is that modeling can be mimicked and an organized scaffold cell matrix with mechanical properties and viable cell density suitable for use as a functional tissue implant can be provided.

  Three-dimensional tissue-like structures such as cell and matrix stratification, alignment, and mesoscale band heterogeneity can also be created that mimic the structure of natural tissues such as tendons, nerves and skin. Nano-microscale cell-independent fabrication of tissue templates with biomimetic structures as described herein is believed to be useful for a variety of therapeutic applications.

One embodiment of the present invention provides a method of making a biomaterial, the method comprising:
Providing a gel comprising a matrix of scaffold fibers and interstitial fluid, and plastically compressing the gel to produce the biomaterial;
Comprising.

  Plastic compression can, for example, increase the density and mechanical strength of the scaffold matrix. The biomaterial produced by the method of the present invention is considered useful, for example, for the production of tissue equivalent implants.

  A tissue equivalent implant is a device for implantation into an individual, for example, to repair or replace internal tissue suspected of being damaged or diseased. Examples of diseased tissue that can be repaired or replaced by tissue equivalent implants include nerves, tendons, cartilage, skin, bones, urogenital elements, liver, cardiopulmonary tissue, kidneys, eye tissue, blood vessels, intestines and glands.

  Preferably, the tissue equivalent implant is a structural biomimetic of the host tissue, in particular that mimics the structure at the cellular level (ie nano-micro (meso) scale). Tissue equivalent implants can contain cells distributed in the appropriate layers, zones, depots and channels, mesoscale, as in natural tissue.

  The structure and distribution of cells in various tissues is well known in the art. Examples of mesoscale layers include epithelial (keratinocytes) and stromal (cutaneous fibroblast) layers in the skin; nerve regeneration duct entrance-, connection- and exit-inclined zones: vascular endothelium (intima), smooth Mention of muscle contraction stroma (2 or more) and outer membrane layer. Examples of channels in an implant can include passages for capillary invasion after nerve transplantation. An example of a zone is a reservoir of hormone-secreting cells held in a construct near the perfused capillary bed.

  In general, suitable scaffold materials can include hydrogels that naturally aggregate or form fibrils, but the fibers of the hydrogel do not have a strong ionic surface charge at physiological pH. The fibers of the scaffold material are preferably not cross-linked or substantially polymerized. Preferred scaffold materials have low hydration / swelling properties to facilitate plastic compression. Many suitable fibril forming scaffold materials are known in the art. A fibril-forming material is a material that includes soluble (insoluble) fibrils and soluble monomer molecules that can self-associate by packing to form fibers.

  Suitable scaffold fibers include naturally occurring macromolecules, eg, proteins such as silk, fibrin, fibronectin, elastin or collagen (eg, type I collagen), glycoproteins such as fibronectin, or polysaccharides such as chitin or cellulose. There is. In certain preferred embodiments, the scaffold fiber is collagen. Collagen types that form natural fibrils include collagen types I, II, III, V, VI, IX and XI, and combinations thereof (eg, I, III, V, or II, IX, XI) ,preferable. For example, type I collagen can be used as a scaffold material.

  In other embodiments, the scaffold may be a synthetic polymer, i.e. a polymer that is not naturally present in the human or animal body. Suitable polymers include organic polymers such as polylactone, polyglycone, and polycaprylolactone, and inorganic polymers such as phosphate glass. Synthetic polymers can be made to reduce the inherent swelling capacity at a selected pH (usually near neutral). Swelling is usually caused by surface charge and osmotic pressure. In such cases, the surface chemical properties of the polymer can be modified to ensure proper hydrophilicity without significant fixed charge density.

  The scaffold may be a composite material comprising two or more different fibers. For example, the scaffold can include fibronectin and collagen, collagen and polylactide, fibrin and collagen, soluble glass fiber and collagen or fibrin, or fibrin and collagen and fibronectin.

  The gels described herein are formed by assembly and stretching of fibrils of the scaffold material, which fibrils form a continuous network around the aqueous interstitial fluid that originally retained the monomer. For example, a triple helical collagen monomer can be first dissolved in dilute acid and then polymerized (aggregated) into fibrils (eg, 37 ° C., neutral pH). As the fibrils polymerize, a phase change occurs and the solid network of fibrils “holds” the remaining interstitial fluid in approximately the same volume and shape—that is, it gels.

  The phase transition from soluble monomer to solid polymer is a characteristic of the gel and is important in providing the properties described herein. The gel is separate from the “sponge”, and the sponge is formed from prepolymerized fibers.

  Unrestricted compression of the gel drains the interstitial fluid and does not revert by removing the load, i.e. the gel undergoes plastic compression. The untreated gel scaffold matrix is generally in a highly hydrated state. Plastic compression of the gel causes the scaffold structure to collapse without loss of structural details, removing moisture from the scaffold in the gel, resulting in increased density and strength.

  The interstitial liquid is typically an aqueous liquid, although other organic solvents may be used in certain abiotic applications. For example, the liquid can be water in which solutes such as salts and proteins are dissolved. In certain embodiments, the interstitial fluid is a cell culture fluid suitable for cell growth and proliferation.

  Preferably, the gel is seeded with cells, in particular human or other mammalian cells. These cells still survive when the gel is compressed into biomaterial.

  The gel may include cells that provide a structure that imparts tissue functionality and replaces or facilitates the repair of body tissue. For example, gels are metabolic cells that provide secretory structures, such as muscle cells that provide contractile structures; blood vessels and / or nerve cells that provide conductive elements; hepatocytes, hormone-producing cells, sebaceous cells, islet cells or adrenocortical cells. Active secretory cells; stem cells such as bone marrow-derived or embryonic stem cells, dermal fibroblasts, skin keratinocytes (and a mixture of the two); Schwann cells for neural implants; smooth muscle cells and endothelium for vascular structures One or more of: cells; urothelial cells and smooth muscle cells for bladder / urethral structures; and bone cells, chondrocytes, and tendon cells for bone and tendon structures. In certain embodiments, the cells seeded into the gel can include fibroblasts.

  Cells can be distributed in any arrangement in the gaps in the gel. For example, the cells may be distributed uniformly throughout the gel, or may be distributed in a limited zone, region or layer within the gel.

  The cells are preferably seeded prior to compression, for example when casting a gel. After mixing the cells with the liquid scaffold matrix, the cells can be seeded into the matrix by solidifying the liquid matrix into a gel. Matrix seeding is performed prior to gel formation under conditions of temperature, pH, ionic strength and shear suitable to maintain viability.

  Depending on the situation, if cells need to be attached to the scaffold, before compression, for 24 hours or less, 12 hours or less, 6 hours or less, 3 hours or less, or 1 hour or less, most preferably from 0 Cells can be incubated in the gel for up to 2 hours.

  The decrease in gel volume caused by compression will be inversely proportional to the increase in cell density in the gel. Thus, the compression of the gel described herein can increase cell density from 5 to 200 times, more preferably from 10 to 100 times.

The initial cell density in the gel can be from about 1 × 10 ⁴ to 1 × 10 ⁷ cells per ml, with about 5 × 10 ⁵ to 1 × 10 ⁶ cells per ml being more preferred. The cell density after compression can be 1 x 10 ⁷ to 1 x 10 ⁹ cells per ml, more preferably 5 x 10 ⁷ to 5 x 10 ⁸ cells per ml.

  The final cell density of the compressed gel can be varied by changing the initial cell density and / or amount of compression, depending on the intended use of the compressed gel. For example, for tissue constructs that require high cell density (such as liver, kidney or gland constructs), the amount of cells seeded into the gel and / or the amount of compression can be increased. The final cell density can be further increased by the secondary compression described herein.

  The proportion of cells in the compressed biomaterial depends in part on the size of the cells. Smaller cells, such as red blood cells, exhibit a lower percentage of compressed biomaterial, even when present at the same cell density, as compared to cells such as larger fibroblasts. In general, the cells represent at least 1% v / v, at least 5% v / v, or at least 10% v / v of the compressed biomaterial. For optimal physical properties, the biomaterial construct can comprise less than 50% v / v, less than 40% v / v, or less than 30% v / v.

  The properties of the biomaterial or implant can be tailored for a particular application or application by changing the proportion of individual components in the biomaterial (eg, expressed as% v / v). For example, changing the ratio of the matrix component to change the strength of the biomaterial, changing the percentage of cells to change the cellular activity of the biomaterial, and / or changing the ratio of microchanneling Perfusion characteristics can be changed.

The plastic compression process can be optimized from a standard starting gel to achieve the desired final ratio of collagen, cells, and channel-forming inserts. A standard gel can contain, for example, 1-4% collagen, 0.2-10 × 10 ⁶ cells / ml, and 0.2-2% channeling fibers or particles.

  Cells in the compressed biomaterial are susceptible to drying. To reduce cell death and / or damage associated with drying, the gel may be compressed in an aqueous liquid, such as a culture medium (such as DMEM, Ham's or Eagle's medium), or a physiological buffer (such as Ringer's or PBS). it can. Any solvent compatible with the scaffold matrix may be used for the cell-free biomaterial.

  Cells are also susceptible to hypoxic cell death and / or damage due to high cell density in the compressed biomaterial. To reduce and / or prevent cell death or damage, the implant or biomaterial can be stored under conditions that maintain viability but do not support cell growth until just prior to use. For example, the implant or biomaterial can be stored at a low temperature, such as 0 to 5 ° C, preferably at 4 ° C.

  The plastic compression described herein is to deform an object such as a gel so that its volume decreases, so that even after the cause of compression is removed, the object is substantially It means a deformation that maintains a new volume. Plastic compression is a rapid cell-independent process that occurs by subjecting the gel to a physical treatment, such as an external force or pressure that drives the interstitial fluid out of the gel. Plastic compression is distinguished from the slow process of cell-driven contraction caused by the intrinsic action of cells growing in the gel, i.e. plastic compression is not mediated by cells, but by the action of cells cultured in the gel. Does not happen.

  Cell-mediated contraction can only cause a small fraction of the compression caused by the plastic compression described herein. Moreover, since cellular contraction occurs with random directional quantities, the overall contraction vector is the average of the entire cell population. Plastic compression has one, two or more constant direction vectors, and the direction, speed, and degree of compression can be controlled.

  Plastic compression makes it possible to accurately define the cell content, matrix density and liquid content of the biomaterial. In addition, one or more of the above parameters in the biomaterial can be locally enhanced, and the spatial position and pattern of such local enhancement can be controlled, eg, stratified, zoned or channeled Bring about

  In certain embodiments, an object that has undergone plastic compression does not show any substantial tendency to swell, but such swelling tendency may later reverse compression without further cell culture or processing steps. .

  In other embodiments, the tissue equivalent implant may undergo some re-swelling during use. For example, a compressed collagen matrix containing chondrocytes does not re-swell initially, but proliferating chondrocytes synthesize proteoglycans / GAGs that take up fluid and swell the implant. Swollen implant structures are particularly useful for mimicking cartilage structures.

  The amount or degree of compression can vary depending on the application of the biomaterial. Gel compression (eg, by compressive force) can result in one or more dimensions of the gel being reduced by at least a factor of 5, at least a factor of 10, or at least a factor of 20. One or more dimensions can be reduced by more than 1/200, 1/150, or 1/100 or more. In a preferred embodiment, the gel thickness is reduced by compression.

  For example, the gel volume can be reduced by 50% or more, 60% or more, 70% or more, 80% or more, 90% or more, 95% or more, 99% or more, or 99.9% or more by plastic compression.

  The time required for compression is less than the time required for cell-driven contraction to occur, but will vary depending on the compression method and conditions used. For example, compression can occur in less than 12 hours, less than 6 hours, less than 3 hours, less than 1 hour, less than 30 minutes, or less than 10 minutes. In certain preferred embodiments, the gel can be compressed in a period of 2 minutes or less, or 1 minute or less.

  Plastic compression of the gel is believed to correlate with losing or removing some or all of the interstitial fluid from the gel.

  For example, the amount of liquid lost or removed from the gel by plastic compression is at least 50%, at least 60%, at least 70%, at least 80%, at least 90%, at least 99% of the original liquid content of the gel, Or at least 99.9%.

  It is preferred that some interstitial liquid remains after compression, for example at least 10%, at least 1%, or at least 0.1% of the original liquid content of the gel. In a preferred embodiment, the dehydration kills the cells and impairs the biomaterial structure so that the gel does not undergo drying or dehydration after plastic compression, such as heat drying, freeze drying, air drying or vacuum drying.

  The degree of compression is determined by the surface charge and hydration equilibrium of the scaffold material. Compression beyond the hydration equilibrium can cause partial re-swelling when the compressed scaffold is placed in an aqueous environment (ie, this compression is partially non-plastic).

  Plastic compression is thought to drain interstitial liquid from the gel or draw interstitial liquid out of the gel. Multiple methods can be used in sequence or simultaneously to compress the gel. In certain preferred embodiments, the compression occurs with one or more directional orientations and the interstitial liquid is discharged from the gel through one or more specific liquid discharge surfaces. The compression vector, and the number and arrangement of the liquid discharge surfaces can be selected to create a specific structure (eg, a lamellar structure) within the biomaterial, as described in the present invention.

  The interstitial fluid can be drained by applying a mechanical force to the gel, for example positive or negative pressure.

  The mechanical force can include a compressive force. The amount of compressive force applied to the gel to achieve the desired compression will depend on the particular environment, but can be readily determined by one skilled in the art. For example, a suitable compressive force can be 0.1 to 10 N, for example 1 N. The gel is preferably unconstrained when subjected to compressive forces.

  Any suitable method of applying a compressive force to the gel can be used.

  For example, the gel can be compressed by one or more of the following methods: applying a static load (eg, dead load) to the gel, applying a load through hydraulic pressure or cams, or passing the gel through a roller.

  In certain embodiments, the gel can be compressed by extrusion through a clamping orifice. The clamping orifice comprises, for example, a conical chamber that compresses the gel as it passes. The conical chamber comprises a porous wall that allows drainage of the interstitial liquid from the gel to facilitate compression.

  In certain embodiments, the compressive force is applied by inflating a balloon within the gel to compress the gel toward the solid surface. The solid surface may form, for example, a tube surrounding the gel, allowing the formation of a tubular implant construct.

  In certain embodiments, the compressive force is applied by centrifugation. A standard hydrated gel is centrifuged at the bottom of the dried centrifuge tube to reduce the liquid content of the construct by plastic compression-the interstitial fluid is drained into the dry porous support or into the collection chamber. The amount of liquid loss can be adjusted by the speed and time of centrifugation.

  Compression methods when the gel is at rest (eg, applying a dead load to the gel or applying a load by hydraulic pressure or cams) are known as “static” compression methods. The method in which the gel moves during compression is known as a “dynamic” compression method (eg, passing the gel through a roller or extruding the gel through a clamping orifice).

  Alternatively and / or in addition to the above, the interstitial fluid can be removed from the gel to cause compression. Any suitable method for removing the liquid can be used. For example, the liquid can be removed by one or more of spill (draining), evaporation, suction, capillary pressure, osmosis, or electroosmosis.

  The liquid can be removed by capillary pressure by “sucking” the gel on the absorbent material. The absorbent material can be paper, in particular blotter paper. This absorbent material can be removed from the gel after compression. For example, a hydrated gel (eg, collagen gel) can be compressed by absorbing it into a porous filter paper sheet through a coarsely woven (> 100 micron) nylon mesh support, but the mass of the material is: It decreases to only a fraction of its original size, and at the same time the density increases. The speed of compression is slower than methods involving mechanical compression. However, the speed of compression can be increased by adding a second mesh blotting layer to the top surface of the gel.

  Removal of the liquid can be done by evaporation, for example by incubating the gel under conditions that promote evaporation (eg, at a pressure below atmospheric pressure and / or at a temperature above room temperature). Since evaporation removes water but not solutes, such methods contain gels as the gels increase in osmolarity as the gel compresses to form a biomaterial. In some cases, it is generally less preferred.

  The liquid may be removed from the gel by aspiration. For example, application of vacuum reduced pressure (eg, mechanical pumps in a multiwell plate system) to a standard hydrated collagen gel placed on a 0.45 micron filtration membrane will determine the size and type of the construct, as well as the degree of vacuum Depending on the liquid content in a few minutes.

  In certain embodiments, the liquid can be conveniently drained by tilting the gel. In that case, the liquid flows out of the gel under the mechanical effects of gravity and gradient. For example, draining a standard hydrated collagen gel with its own weight in an angled direction in a room temperature humidification chamber will provide adequate liquid loss and compression in 3 hours. In three days, exponential abrupt liquid loss occurs, reducing the gel mass by a factor of six.

  Certain plastic compression methods, such as osmosis, electroosmosis, or evaporation, specifically remove the solvent from the gel but not the solute. This can change the ionic properties (eg, salt concentration, pH) of the liquid in the gel.

  When seeding cells in a scaffold gel, the gel environment is preferably maintained at physiological conditions (eg, temperature, pH, moisture, and ionic strength) so that the cells can survive. In such biological embodiments, it is preferred that plastic compression does not significantly change the ionic properties of the gel liquid from physiological conditions.

  In an abiotic embodiment, if the gel does not contain cells, the gel environment need not be physiological, including methods of changing the ionic properties of the gel liquid, such as osmosis. Any method is suitable.

  Two or more different compression methods can be used sequentially or simultaneously to compress the gel. For example, the gel can be plastically compressed by mechanical compression, but preferably at the same time the liquid can be removed from the gel by one or more of the following methods: suction, capillary pressure, osmosis, electroosmosis. In some embodiments, a combination of capillary pressure and mechanical compression is used for rapid compression.

  The methods described herein can include more than one, eg, 2, 3, 4 or more distinct stages of plastic compression.

  As described above, interstitial liquid can exit the gel through the limited liquid discharge surface of the gel. In certain embodiments, for example during plastic compression of a collagen gel, the outer “skin” (liquid drainage surface) of the gel can function as a filter and allow liquids and solutes to exit, but collagen fibrils, cells, and / or Or other macromolecular aggregates such as particle inclusions, growth factor depots / vesicles, and mineralized particles (eg, bone mineral particles or ceramic biomimetic particles) can remain on the gel surface. In addition or alternatively, the gel can be supported on a porous sheet or membrane that assists in filtering the liquid to be removed from the gel. This can cause some local non-uniformity of compression on the gel surface (especially with certain liquid “exit” surfaces). Thus, a dense layer of cells, collagen or other macromolecular aggregates can be localized on the gel surface. This lamellar structure is thought to be useful for making tissue equivalent implants. The scale of such fabrication techniques can be adapted, for example, to create structural features or structures at the nano or micro level.

  The properties of the scaffold matrix are important in determining how tissue repair or regeneration is organized by the cells, and the initial scaffold structure effectively determines the majority of the subsequent three-dimensional structure. Thus, in many applications it is useful to align cells and fibers in tissue equivalent implants.

  The cells and / or fibers of the scaffold matrix can be aligned, for example, by applying tension across the gel. Tension can be applied before, after, and / or during plastic compression.

  The tension is preferably uniaxial and can apply 5-50% uniaxial strain, preferably 10-30% uniaxial strain to the gel. The fibers and, if present, the seeded cells are aligned in an orientation parallel to the direction of the principal strain.

  For example, for collagen gels, 5-30% strain, preferably 20-25% strain can be utilized.

  The methods described herein can include tensioning the gel to align the scaffold material fibers and / or cells.

  For example, if the gel is placed under a uniaxial tensile load at the beginning or midway of the plastic compression process (eg, 20-60% compression) and then returned to the plastic compression plan, the fibers induced by tensile strain Is aligned in the gel, resulting in an aligned dense composite material, which in some embodiments may include living cells.

  Alternatively, the fibers can be aligned by clamping the ends of the gel and applying a uniaxial strain across the major axis prior to plastic compression of the gel.

  The alignment of cells and fibers within the scaffold matrix is described in Eastwood et al. (1998).

  In addition to the scaffold, cells, and interstitial fluid, the gel may further include other components. In particular, the gel can comprise solid elements such as capillary filaments or porous beads.

  Capillary filaments are rigid solid polymer insoluble or soluble fibers. Suitable filament diameters are preferably less than about 100 μm.

  Soluble filaments added into the gel can dissolve to form capillary channels within the gel. Such capillary channels within the gel may be useful, for example, for one or more of the following purposes: perfusion, drug and / or gene and / or medium delivery into the scaffold; and recipient's Merge with the circulatory system.

  Suitable soluble filaments can be made from soluble phosphate glass, polycaprylolactone, polyacetate, polyglycolic acid, silk, polysaccharides, or molten or crystalline salts.

  Insoluble filaments are considered useful for optical therapy, optical monitoring, signal transduction, and / or strain detection. Suitable insoluble filaments are made from glass.

  Capillary filaments can be placed between gel layers, added to the gel before casting, or placed in the gel after casting. In embodiments that include two or more cast gels, the filament may be sandwiched, for example, between the gels.

  In certain embodiments, the gel can include porous beads. Plastic compression of the gel matrix pushes the scaffold fibers into the pores of the porous beads, giving a tightly bonded structure. This is considered useful as a substitute for artificial bone or calcified cartilage, for example, when seeding osteoblasts or chondrocytes.

  Suitable porous beads are about 100-500 microns in diameter, for example, solid materials such as porous ceramics, glass, phosphate glass, hydroxyapatite, or bone mineral preparations (removing organic phase from natural bone) Can be made from

  The particle: gel: cell ratio depends on the size of the particles and the desired tissue properties (eg, tightly or loosely packed hard tissue).

  In certain embodiments, a linear construct (eg, a rod, strip, or roll construct) can include the porous beads described herein at both ends, eg, by screwing directly into bone in vivo. It is possible to facilitate the fixation of the construct in

  Multilayer biomaterials are useful. This is because natural tissues are layered at the micron level, and these layers are thought to play some role in the organization of cellular tissues and the separation of cellular functions.

  The methods described herein allow for the production of biomaterials with a fine structure, eg, biomaterials having multiple layers (ie, two or more layers). Such biomaterials are particularly useful in making tissue equivalent implants.

  Layers also facilitate tissue movement (layers that respond to gliding, flexion), help perfusion through the layers, and (for example, in tendons) rapid anisotropic nutrient transfer that moves between layers along the layers And direct ingrowth of capillaries, nerves, stroma, etc. after transplantation (inner growth in the least resistant plane or channel).

The method of making a biomaterial comprises:
Overlaying the first and second gels, each gel comprising a scaffold fiber matrix and interstitial fluid, and plastically compressing the gel to produce a biomaterial (the first and second gels) The second gel forms a layer in the biomaterial).

  In certain embodiments, the first and second gels may be different regions of the same gel, and these regions are brought into contact, for example, by folding or rolling the gel.

  In other embodiments, the first gel and the second gel may be different gels, i.e., two or more gels may be stacked.

  The two or more gels may have different functions and / or may differ in terms of one or more physical properties. For example, gels may differ in one or more of the following: fiber organization in the scaffold, fiber type, scaffold density, and macromolecular aggregates present in the gel, porous Concentration, type or mixture of mass particles and / or biological depots (eg, for growth factors, hormones, enzymes and / or minor matrix proteins).

  The gel can be plastically compressed to produce a single sheet of biomaterial having two or more layers.

  Preferably, the gels adhere well so that they do not easily separate under mechanical loads, for example, the interface is preferably 1-10% of the fracture strain of the layer itself, preferably May be stable for 5-10%.

  When two or more gels are combined, complex tissues and organs can be built up one by one. This may be useful, for example, in making implants to repair or replace nerves, tendons, cartilage, skin, bones, genitourinary elements, liver, cardiopulmonary tissue, kidney, eye tissue, blood vessels, intestines and glands .

  For example, an implant for repairing neural tissue (PNS or CNS) having cells around a collagen outer skin and a core of a nerve guide material (eg, fibronectin) can be made. Cell / growth factor (eg, NGF) depots may be placed at the inlet and outlet ports.

  In order to produce an implant for repairing a ligament, for example, a ligament implant having a hard bone attachment device at each end for incorporating hydroxyapatite beads or particles into the end of a PC collagen rod and fixing it in place. To form.

  In order to produce an implant for cartilage repair, for example, a collagen roll-like or helical structure seeded with articular chondrocytes (illustrated) is produced, and secondary compression is applied to this roll (with respect to the long axis). (Vertical), incubate in standard growth medium for about 7 days. Endogenous chondrocytes re-swell the construct (by proteoglycan synthesis and water uptake) to form a cartilage implant.

  A connective tissue (or other tissue) implant with a given vascular channel uses collagen to form the above helical (or layered) construct, but parallel fibers made of phosphate glass (ideally Can be made by placing between 20 and 200 microns in diameter) between the layers. When the glass fiber dissolves in an aqueous medium (eg after implantation or in a bioreactor), a channel remains in the collagen implant that serves as a guide passage for the ingrowth of capillaries (nerves). . Such a structure follows and fills the above channels as the path of least resistance.

The methods described herein can be used, for example, as part of an integrated, automated culture process, with appropriate physical loading on the gel (e.g., preferably as part of a series of pre-programmed events). It can be carried out well with a mechanical bioreactor suitable for applying a certain compression and tension. The mechanical bioreactor is preferably suitable to allow culturing and / or incubation of three-dimensional tissue equivalent implants, for example, uniaxial tension means (eg, subjecting the construct to periodic or inclined tensile loads) ), And optical and electrochemical (O ₂ and glucose) sensors.

  Alternatively, tissue equivalent implants (cellular or abiotic) can be made using a continuous flow-through method (conveyor belt type) having a series of steps including a plastic compression step.

One aspect of the present invention provides a method of making a tissue equivalent implant from a biomaterial made as described above. For example, a method for making a tissue equivalent implant is:
a) plastically compressing a gel comprising a scaffold fiber matrix and interstitial fluid to produce a biomaterial, or
b) superimposing two or more gels, each gel comprising a matrix of scaffold fibers and interstitial fluid, and plastically compressing these gels to produce a biomaterial having two or more layers;
Either
Making the implant from the biomaterial;
Comprising.

  Each gel can further include living cells as described above.

  Plastic compression can be performed as described above. In certain preferred embodiments, the gel can be subjected to mechanical compression.

  Tension can be introduced across the gel to align cells and scaffold fibers. Tension can be applied before, after, or during plastic compression of the gel.

  As described above, the method of the invention is performed at a temperature below room temperature (eg, about 4 ° C.) so that the seeded cells still survive despite the increased perfusion distance as the cell density increases. be able to. Additional gel layers may be added as described herein.

  The cell density and mechanical properties of the implant depend on the conditions of plastic compression.

  Biomaterial can be shaped and / or shaped to produce the implant. The gel can be formed into a predetermined shape, for example during casting and / or plastic compression. In certain embodiments, this shape can be further shaped after the plastic compression step and / or subjected to additional compression.

  Tissue equivalent implants can be molded or shaped into accessible implant shapes, such as patches, tubes, tapes, strips, rings, rolls, sheets or threads. The final shape of the tissue equivalent implant depends on the particular situation in which it is to be used.

  The tissue equivalent implant can be formed into a pliable shape suitable for further shaping. Plastic compression can be, for example, symmetric or asymmetric.

  After the implant construct is manufactured by plastic compression, the construct may be immediately implanted into an individual, but can be further incubated to adjust the subtle characteristics of the scaffold matrix, such as stiffness, elastic modulus, and the like.

  In an embodiment in which a biomaterial sheet is produced by plastic compression, the biomaterial sheet can be rolled and formed to form a roll of biomaterial.

  This allows a three-dimensional implant to be rapidly assembled from one or more flat two-dimensional biomaterial layers, but such implants can be handled well by the surgeon.

  In other embodiments of producing a sheet of biomaterial by plastic compression, the biomaterial can be folded to produce a construct comprising two or more layers. This construct can be further cut or shaped as required.

  The construct can be wrapped or folded around a central fiber / needle to create an implant with a central channel.

  Soluble fiber (eg, phosphate glass) can be placed between the layers of the construct. These fibers then dissolve and leave channels for in vivo angiogenesis or nerve reinnervation.

The construct is wrapped or folded around a sensor / monitor fiber (eg, 100-200 micron optical fiber for O ₂ monitor), for example, after implantation or in a bioreactor, continuously in a cellular matrix structure state Real-time measurements can be obtained.

  Rolled or folded multi-layer biomaterials may be used as tissue equivalent implants as is, but further plastic compression can be used to adhere layers together, achieve desired dimensions, cell density, etc. And other properties can be improved.

  Secondary compression dramatically improves biomaterial properties. For example, once a multi-layer construct comprising live cells is formed, it can be subjected to secondary compression, draining more interstitial fluid and providing higher strength. Secondary compaction of the roll-like construct typically removes 60% of the remaining liquid and gives a dense strap of strong material-can increase the fracture stress by more than 3 times (0.6 megapascals to 2 Megapascal). Secondary compression can be performed using the techniques described above.

  Rolled or folded constructs can be stored under conditions that maintain cell viability but reduce or prevent cell proliferation. For example, the roll can be stored at 0-5 ° C. Such conditions prevent cell death or damage associated with high cell density, such as cell damage due to hypoxia.

  If the biomaterial has one or more layers or regions, the layers or regions can be arranged to provide a complex roll or spiral assembly of separate zones or regions, It may be useful for certain implant applications. For example, the layers can be concentric with respect to the outer layer and one or more inner layers. In addition, various construction shapes can be obtained by rolling up the sheet in various directions.

  For example, a rolled implant construct comprising four separate regions for use in neural tissue repair and / or replacement is shown in FIG. This construct is composed of an outer sheet 1, a nerve entry substrate 2, a long-distance fiber guidance material 3, and a nerve exit material 4 containing growth factors, and after compressing these layers, it is rolled and rolled to the final Made by forming a helical assembly.

  A biomaterial that is folded or rolled can be used as a tissue equivalent implant. That is, it can be introduced into the human or animal body for repair or replacement of the underlying tissue.

Another aspect of the invention provides a method of treating an individual's damaged tissue, the method comprising:
Creating a tissue equivalent implant using the methods described herein, and fixing the implant to the damaged tissue to repair and / or replace the tissue;
Comprising.

  The implant can be secured by any convenient technique. For example, it can be sutured or glued in place.

  Any of the constructs described herein (eg, rolled, folded, secondary compressed biomaterial) can be sutured and applied to a body part even when performed under muscle load. Can be sutured surgically.

  Another aspect of the invention provides a biomaterial or tissue equivalent implant made by the methods described herein.

  Aspects of the invention will now be described by way of example and not limitation with reference to the accompanying drawings and the following examples. Further aspects and embodiments will be apparent to those skilled in the art.

  All documents cited in this specification are hereby incorporated by reference.

  Table 1 shows the average of the collagen construct thickness (monolayer) at three stages of compression.

  Table 2 shows the cell viability at various stages during the compression process with tension or drying applied from the starting cell preparation and freshly cast gel (before compression) to after compression.

Method
Cellular and acellular collagen gel formation As described previously [1, 21], acellular collagen gel was prepared using 2.4 mL of sterile rat tail type I collagen (2.16 mg / mL protein in 0.6% acetic acid: First Link (UK) Ltd, West Midlands, UK), 0.3 mL of 10-fold concentrated eagle minimal essential medium (EMEM) (Gibco Chemicals, Invitrogen, Scotland, UK) and 0.3 mL of eagle balanced salt solution (EBSS) (Sigma-Aldrich Co , Dorset, UK) was prepared by neutralization with 5 M sodium hydroxide (Merck, Leicestershire, UK). The cells were seeded gel, 10% fetal calf serum (FCS) 0.3 mL of Dulbecco's Modified Eagle Medium 10 ⁶ human dermal fibroblasts suspended in (DMEM) containing, in place of EBSS, neutralizing Immediately afterwards, it was mixed with a collagen solution. The cell density at the time of casting was 0.33 × 10 ⁶ cells / ml collagen gel. Pour the gel into a rectangular mold (33 x 13 x 4 mm) and attach in a 37 ° C CO ₂ incubator for 70 minutes for cell-free gel (experimental range 35-180 minutes), for cells attached to cellular gel Solidified / stabilized for 1-3 hours. The measurement of liquid loss (self-compression) from the collagen gel was based on the% wet weight loss. Transfer of the gel from the casting chamber for the first wet weight measurement uses a custom flat scoop spatula to minimize manipulation, then incubate in a sealed humidified chamber (room temperature), 0.5, The weight was re-measured after 1, 2, 3, 4, 24, 48, 120 hours (horizontal or 45 ° tilt: n = 3 for each measurement). Since almost all of the wet weight change occurs within the first 3 hours, a horizontal 3 hour “self-compressing” gel was used as a stable control (giving 57% compression).

Cell culture and cell viability assays As previously described [21,22], exfoliated (newly obtained from the operating room with sufficient ethical approval immediately after surgery for breast reduction or abdominal wall formation) Normal human skin fibroblasts were prepared by cultivating the defatted human skin in a normal graft culture. Fibroblasts growing from 2-4 mm adherent tissue fragments were treated with 10% FCS (First Link (UK) Ltd.) and streptomycin / penicillin (500 μg / ml and 500 respectively) under conditions of 5% CO ₂ and 37 ° C. Grown in Dulbecco's Modified Eagle Medium (DMEM: Gibco, Paisley, Scotland) containing ICN Biochemicals, Thyne, UK. Cell viability measurements consisted of constructs in DMEM + 10% FCS containing 1 μM calcein AM (detecting living cells) and 1 μM ethidium homodimer (staining dead cells) at each stage of plastic compression. This was done by incubating for 1 hour. Live and dead cells appear green and red on the fluorescence microscope, respectively, and the survival: kill ratio is the cell count from the micrograph through the grid (total 4000 cells counted from a random area) Estimated by

After plastic compression and tension pre-alignment solidification and incubation, the gel was compressed as usual using a combination of compression and blotting with layers of mesh and paper sheets, as shown in FIG. Briefly, a 165 μm thick stainless steel mesh (mesh size about 300 μm) and a layer of nylon mesh (mesh size about 50 μm), two layers of water absorbent paper with two stainless steel spacers (thickness 300 μm, Figure 1a) Placed on top. Place the collagen gel on the nylon mesh (between the spacers), cover with another nylon mesh, load with 50g flat plastic block (Delrin polymer) for 5 minutes at room temperature, 2 nylon meshes A flat collagen sheet (thickness 20 to 40 μm) protected between the two was obtained. Characteristic evaluation of plastic compression only (ie, no enhanced blotting) was without water absorbent paper.

  The blotting inevitably contributed slightly to the final compression, because the applied load was limited by a 330 μm spacer shim. The weight of the gel was measured before and after compression to obtain% compression. The load and duration were 10-50 g and 10 seconds-5 minutes. The thickness of the construct after compression was measured directly from cross-sectional histology specimens (stabilization control: 57%, compared to gels compressed to 87% and 91% by weight). For “tension pre-alignment”, the compressed collagen sheet (with or without cells) is clamped horizontally in the culture chamber (6-7 mm overlap at both ends) and the thread is moved for approximately 1 minute. Used to strain up to 25%. The deformed gel was fixed in 0.1 M sodium cacodylate buffer pH 7.4 containing 4% paraformaldehyde for microscopy.

Helical assembly of collagen sheets and composites
The PC collagen sheet was wound along the minor axis immediately after compression and peeled off from the nylon mesh to produce a rod-like product (diameter 13 mm x 1.75 mm) that was tightly wound in a spiral shape (Fig. 2a). Collagen: Fibronectin (Fn) and Collagen: Hydroxyapatite (HA) were used as examples of heterogeneous helical assemblies. As previously reported [23], an oriented fibronectin mat was prepared in a stirred ultrafiltration cell. Briefly, plasma fibronectin [19] (Fn: 1-2 mg / mL in 2M urea, Bio-Products Lab, Elstree UK) was concentrated (10 kDa molecular weight) in a stirred ultrafiltration cell at 4 ° C. Cut membranes-Millipore, Watford, UK). The resulting Fn mat was rinsed, lyophilized, and rehydrated immediately prior to co-compression. The collagen gel was displaced (partially covered) from the Fn mat and placed on top for compression. The helical assembly of this composite was the same as before, starting from the offset Fn end. Collagen: For the HA example, approximately 200 mg of HA particles (106-180 μm, provided by Plasma Biotal Ltd., Derbyshire, UK) were spread evenly over the long edge surface of the newly cast collagen gel (5 mm width) Until). The compression of these types of constructs was done by simply sucking the liquid into the underlying paper layer without using any upper nylon mesh or load. The helical assembly of this construct (wound around the short side) gave a dense collagen cable with a hard end that was embedded and compressed with HA.

Mechanical Testing Mechanical testing of single PC collagen sheets is not reliable due to the thin structure (20-40 μm), but the helically assembled construct was used throughout. Dynamic mechanical analyser (DMA-7e: Perkin Elmer, Buckinghamshire, UK) was clamped at the ends of the structure stored in EBSS prior to testing with a thin steel mesh 2mm strip reinforced with cyanoacrylate strong adhesive Attached to. The mechanical test was carried out under uniaxial tension at a loading rate of 200 mN / min (stress rate of 83 KPa / min) and constantly hydrating the sample by use of a buffer until failure. Quasi-dynamic stress-strain values for each mechanical test were obtained with Pyris ^™ version 5.02 software (Perkin Elmer, Buckinghamshire, UK) connected to DMA-7e. Statistical analysis was performed with Microsoft Excel ^™ software.

Samples for imaging and image analysis light microscopy are fixed with 0.1 M sodium cacodylate buffer pH 7.4 containing 4% paraformaldehyde and embedded in wax as usual to produce sections (5-8 μm) (collagen Stained with hematoxylin / eosin (H & E) or Picro-Sirius Red (to increase birefringence of the fibrils). The specimens were examined with an Olympus BH-2 micrograph. PC construct (93% compression) for transmission electron microscopy (TEM) examination aimed at analysis of collagen fibril density, fixed in 4% paraformaldehyde, 0.1 M sodium cacodylate buffer pH 7.4 and 2% glutaraldehyde, After rinsing in cacodylate buffer, it was treated with 1% (w / v) osmium tetroxide, dehydrated as usual, and embedded in Spurr resin (Agar Scientific, Stansted, UK) to prepare ultrathin sections. Stained sections (2% (w / v) uranyl nitrate and Reynold lead citrate) were examined on a Philips CM12 instrument. Random images (same magnification) obtained from the edge of the PC construct replica and the gel body region were digitized to outline and measure the total field area% occupied by collagen fibrils (ie, measurement of fibril density) ). For analysis of captured images, Openlab ^™ image analysis software, version 3.1.5 (Improvision®, Coventry, UK) was used. Graphpad Prism ^™ software, version 4 (GraphPad Software®, San Diego, USA) was used for statistical analysis. For scanning electron microscopy (SEM) examination, the PC constructs were fixed in 4% paraformaldehyde (0.1 M sodium cacodylate buffer pH 7.4; 2 hours) and then 1% tannic acid (w / v) After being fixed (0.05 hour) in 0.05 M sodium cacodylate buffer containing water, it was dehydrated through an alcohol series, treated with hexamethyldisilazane (HMDS) and air-dried. The dried specimen was cut along one edge to reveal the internal structure, sputter coated with gold palladium and examined with a Joel 5500LV SEM (L. Wollweber et al., J. Microsc. 1981, 121, 185 ).

result
Collagen gels embedded with plastically compressed cells (1 million in total) were compressed using the system shown in FIGS. 1 and 2 by dead load (50 g) and removal of the effluent by perforated paper. The casting gel had a width of 13 mm, a length of 33 mm, and a depth of about 4 mm. The compression was almost the same at 1 minute and 5 minutes, but a standard compression time of 5 minutes was used. The “stop” shim was designed to leave a 300 μm thick (ie, about 10-fold compression) PC construct, but the process was exceeded due to the capillary action of the large perforated paper sheet, resulting in a final The sheet-like construct was found to be 30-50 μm thick at the time of measurement. This corresponds to> 100 times compression as the reduction ratio (1.72 ml to 0.0172 ml).

The initial concentration of cells in the gel was approximately 5.8 × 10 ⁵ cells / ml, which increased to 5.8 × 10 ⁷ cells / ml after compression. If the cells were spherical with a radius of 10 μm, the cells accounted for approximately 0.18% v / v of the starting gel and 18% v / v of the final sheet-like construct. Of course, the larger the cells at the same density, the greater the proportion of the construct, the smaller the cells.

At the end of the composition compression step of the construct (5 min; 50 g), the cell density was approximately 15-20% of the total construct by volume (assuming a spherical nominal volume of cells). At this stage, the total liquid loss was measured as 98.4% by weight and the wet weight of the sheet-like construct was 43.5 mg. The weight of the gel air-dried at 37 ° C. for 24 hours was 5.4 mg, indicating that this sheet-like construct contained 18% collagen and 82% liquid. A basic composition consisting of 15-20% v / v cells (assuming spherical cells with a radius of 10 μm), about 20% collagen, and the balance water, is comparable to some simple natural tissues. In the case of spherical cells with a radius of 5 μm, the cell content is reduced to 2% v / v.

The cell viability of the construct and a construct comprising live rat tendon fibroblasts in cultured collagen gel was assembled and compressed as described above.

  Fibroblast viability was accurately measured immediately after compression (at a sheet thickness of 40 μm). A fluorescent double staining method was used, in which the construct culture was treated with 1 μM ethidium homodimer and 1 μM calcein AM for 1 hour at 37 ° C. in DMEM medium. As a result, when viewed with a fluorescence microscope, the red fluorescent dye is deposited in the nuclei of dead cells, and the cytoplasm of living cells becomes green fluorescent (respectively). Random fields of view were photographed over the entire construct (each containing a total of 2 to 300 cells) (in all cases under a 40x objective), and red and green fluorescent cells were taken for each of the photographs Counts were made using a 25 mm grid in the field of view. The ratio of live / dead cells was measured from approximately 4000 cells from 15-20 fields per measurement and expressed as mean survival rate +/− SD (Table 2).

  The gelation process itself (complete hydration) had no significant effect on cell viability and almost no effect on cell viability (decreased by about 10%). Even with the relatively high compression / shrinkage of the constructs used herein (decrease by a factor of 100 in 5 minutes), the survival rate is only slightly reduced (about 6%) during plastic compression, and the tensile reserve There was no reduction due to the alignment step. However, the compressed construct is in the form of a sheet with a thickness of about 40 μm, and as a result is very vulnerable to dehydration. Table 2 shows the rapid decrease in cell viability observed due to dehydration, simply by keeping room temperature air flowing in a sterile flow hood (20 seconds and 1 minute, respectively). This resulted in about 70% cell death in 20 seconds, and viability was completely lost after 1 minute (ie, total dryness). At this stage, cell death could be reduced with procedures performed either in liquid (medium / PBS) or in a 100% humidified chamber.

  Note that in experiments where a more controlled thickness was required, blotting paper was removed and the method was performed using only the dead load while submerged in the medium.

Gel Aging Uncompressed Collagen Gels were considered “unstable materials” in any form of lifting / manipulation that lost liquid and changed its density / properties outside of an aqueous environment. The instability was examined in terms of the initial liquid reduction rate by simply leaving the gel on a nonporous plate that was flat in a humidified (100%) chamber or inclined at 45 ° C.

  FIG. 3 shows the decrease in mass (liquid) over time for the gel after the incubation period that had been increased at 37 ° C. (collagen gel continues to stabilize for about 2 hours after initial solidification). The sudden drop in mass during the first 2 hours indicated that it was essentially mechanically unstable once removed from the liquid. Self-compression suddenly decayed to a slow compression rate with a hydration level that correlated with the length of the stabilization period.

  Increasing the flow gradient by tilting the gel angle to 45 ° C during self-compression increased the duration of the initial rapid compression (Figure 3) and greatly increased the level of plastic compression. A 180 minute self-compression resulted in a mass loss of approximately 57%, but produced a “control” gel stabilized by this compression.

Dehydration of the construct in air / water loss Cell-free collagen construct (in this case, compressed 40 μm sheet before rolling) is prepared fresh as above (average 3) and placed on a precision balance The rate of water loss due to evaporation was measured by standing in the air (no airflow and room temperature) and measuring the weight loss (FIG. 4). The weight loss was linear in the first stage and the inflection point of the curve (about 42 minutes) showed that almost complete dehydration was reached after only 40 minutes (no desiccant in still air) Note that point). This indicates the expected time for air drying.

Alignment in compressed construct Standard hydrated collagen gel was prepared as described above and subjected to plastic compression.

  In the middle of the plastic compression process, the gel was placed under uniaxial tensile load (25% strain) and then returned to plastic compression. The alignment of fibrils caused by tensile strain was fixed in the gel by plastic compression, resulting in an aligned dense composite. Cells seeded in the gel at the start of the process were alive in the finished composite.

Cell-free collagen PC construct (sheet-like) after pulling and pre-aligning collagen fibrils by applying a uniaxial tensile load that gives an applied strain of 25% as a whole for tensile pre-aligned collagen PC sheet scanning electron microscopy ) Was examined using a scanning electron microscope. The specimens were fixed in glutaraldehyde as in the case of normal SEM, placed on a SEM pedestal, dried at the critical point, sputter coated with gold, and then examined with a JOEL semi-environmentally controlled SEM.

  FIG. 5 (upper figure) shows the surface topography after compression with a tightly raised pattern of a nylon support mesh (having a fiber structure of <1 to> 100 μm). The material surface was observed to consist of parallel thick collagen fibers (often increased beyond the starting collagen fibril diameter of 50-100 nm). When 25% uniaxial strain was applied, the raised pattern realigned, resulting in surface grooves parallel to the applied strain. Since this process is a plastic deformation, the surface restructuring is plastic, resulting in a stable engraved topography controlled by a combination of relief and tension pre-alignment. . In conclusion, this μm and sub-μm anisotropy pattern is stable after the applied tension is removed, and a directional lamellar structure containing compressed collagen fibrils was also seen (FIG. 5, bottom). Figure).

Effect of blotting method The relative importance of liquid uptake (sucking) by two porous layers (paper and nylon) was investigated.

  A rectangular strip of two materials was held vertically in a PBS solution containing blue pigment (to facilitate visualization), and the height at which the top of each liquid rose was measured over time. . The results are shown in FIG.

  Although there was a difference in the pattern of liquid uptake by the two materials, both could significantly contribute to the blotting effect.

  In one experiment in this study, the paper layer was removed to reduce the contribution of blotting to overall liquid removal. This reduces rather than invalidating the blotting effect. Therefore, the compression performed using the lower layer of paper was called the enhanced blotting method. Non-reinforced compression was performed using only the nylon layer.

  For example, compression performed while submerged in the liquid to reduce drying completely eliminated the blotting effect.

Compression conditions The replicated collagen gel was compressed (without reinforced blotting) without blotting the porous paper layer under various dead load and compression time conditions. In each case, the starting and final (compressed) weights of the construct were measured to obtain a% gel weight reduction, and the results are shown in FIG.

  The average compression with 10 g load and 10 seconds was 57%, but increased to 88% when 50 g, 1 minute was applied (FIG. 7). After 2 minutes using a 50 g compression load, little additional compression was observed, and extending the compression time to 5 minutes did not increase liquid loss, but decreased variability between constructs. A 50 g, 5 minute vertical compression protocol was adopted as the standardized method. This PC protocol reduced the thickness of the initial construct from a collagen gel of approximately 3.6 mm to a 23-48 μm collagen sheet (Table 1). That is, the shrinkage was> 2 digits. Such collagen sheets were typically 18% collagen and 82% liquid. Collagen concentration was increased for the cell seeded constructs due to a decrease in the relative liquid content and the final cell density was 10-20% (v / v), which was initially seeded. Depending on the density and the diameter of the non-adherent (spherical) cells. Such a composition would be comparable to some immature connective tissue and would be suitable as an early tissue graft.

  Complex effects were analyzed by comparing the results of 2 minute and 5 minute compressions with and without blotting enhancement. Even with enhanced wicking, no additional compression effect was observed at either 2 minutes or 5 minutes compression.

The collagen fibril density of a collagen fibril density standard construct (40 μm thick compressed sheet: 50 g, 5 minutes) was measured relative to the control construct, which was under gravity on a plate in a humidified chamber. Stabilized to about 50% of the original weight by leaving to compress.

  The fibril density was directly measured by image analysis on the ratio (%) of collagen fibril area to total screen area obtained from TEM images (all at the same magnification).

  Collagen specimens were prepared with a normal transmission electron microscope, fixed with glutaraldehyde (2.5% in cacodylate buffer pH 7.5), embedded in Spurr resin, and ultrathin sections (thickness) And stained with lead citrate as usual, and then examined with a Philips CM12 microscope. High quality photomicrographs were taken from random good quality fields from the test gel (“center” fields were taken randomly from both sides of the midline of the construct for a center-edge comparison). These images were then loaded into an image analysis system (Improvision-Mackintosh, Birmingham, UK), where the replica collagen fibrils were outlined in the replica field of view to distinguish the fibril area from the space between the fibrils. . This image analysis method then yielded a quantitative estimate of the fibril density (%) per field. Median density is shown with standard deviation and range (box mustache). Comparison of the density of the control and compression constructs shows a dramatic increase in fibril density.

  Collagen fibril density in the central and edge regions of the compressed construct was compared both with and without cells (Figure 8). Collagen fibril density was high in the cellular constructs in both the marginal and central regions, as expected because the volume occupied by the cells was excluded by the measurement (this was Effectively replaced the collagen content). The increase was about 50% at the edge and 100% in the central region, in both cases p <0.002. The fibril density in the edge region of the cell-free construct was about 50% higher than the corresponding central region (p = 0.0038). However, this difference disappeared completely in the cell seeded construct (obviously due to the sampling technique and the local separation of the cells). Separation of collagen fibrils towards the edges of the construct (ie, lamella formation or heterogeneity), as expected (above), is effective filtration behavior of the liquid from the collagen gel surface during compression, as well as Occurs based on complete retention of collagen fibrils at the interface.

Tubular construct highly hydrated material (gel) —An alternative to compressing the cell construct is to generate a compression impulse from the core of the structure (FIG. 9). This may be useful, for example, to form a tube with a thin tube wall. A material (gel) device is formed to cover the outer surface of the cylindrical balloon, but by expanding the core balloon, the construct is expanded (applying the desired mechanical load if necessary) and porous (generally cylindrical) The outer surface of the construct is pressed against the mold (of shape), thereby achieving controlled liquid ejection. This is due to the size of the layer being made (eg wide aperture, thin walls), the laminate structure, the “out” direction of the liquid (ie the microflow vector), and the cells and fibers present in the support. This is advantageous in terms of applying mechanical force (eg for alignment).

Continuous Flow Extrusion-Plastic Compression Continuous Flow Extrusion-Plastic Compression can be achieved by liquid extrusion, tissue shrinkage / compression through a conical flow chamber (FIG. 10). Treat the pregelled sol of a (collagen) dope preparation (including cells if necessary), which undergoes a directional increase in shear force, an increase in concentration, and an increase in fluid viscosity simultaneously Like that.

  This can be achieved by taking advantage of a few minutes of gel time (the gel time can be controlled and extended by holding the temperature at 4 ° C and raising it to 37 ° C). . Prepare a pre-gelling solution of collagen (or other similar liquid crystal polymer such as fibronectin, silk protein or fibrinogen) and slowly pump it along the narrowing conical chamber, but the chamber wall It is porous and enables the extruding of interstitial liquid. As the gel-sol liquid moves along the chamber, it undergoes a directional shear force (aligns and assembles the polymer fibers); the liquid is pushed out of the dope into the porous wall (the polymer polymer covers the inside) Held by the semipermeable membrane), the material is compressed as it gels; the sol quickly solidifies into a gel, and the gel is pushed out of the clamping orifice at the tip of the conical chamber so it is rapidly compressed and solidified To do. The result is an extruded dense cable or thread of aligned “tissue”.

Three constructs ( multilayered construct fibroblasts) (flat rectangular collagen gel) were contracted for 16 hours (non-external) as described in Mudera et al., 2000, and Prajapati et al., 2000 in a uniaxially constrained arrangement. ), Alignment of cells and collagen fibrils was achieved. At zero time, each construct (or “leaflet”) was approximately 3 mm thick, with a total thickness of 9 mm. At the end of the contraction culture, the thickness of each leaflet was 2 mm (6 mm in total).

The three leaflets were released from restraint and placed in a 35 mm culture dish supported by (1) coarse nylon mesh, (2) a sterile filtration membrane with a pore size of 70 μm. The leaflets were then covered with DMEM medium (Dulbecco's modified Eagle medium) and placed under a compression platter in a Dartech mechanical test apparatus and placed in a 37 ° C. CO ₂ cell culture incubator.

  The triple leaflet construct was maintained in culture for 5 days so that the construct was 1.5 mm thick overall (ie, the overall density increase was 6-fold from the fibril alignment and plastic compression stages) ). The resulting construct was (optically) very dense and stiff, and each of the three leaflets adhered firmly to the adjacent leaflet, resulting in a fully integrated single construct.

  This indicates a surprisingly effective mechanism of plastic compression, removing moisture from the collagen fibril construct and integrating the bioreactor sheet at the end of the culture period (giving the desired thickness and physical strength). The resulting construct did not return to its original size (ie, the density did not decrease), contained live cells, was mechanically integrated and strong.

Capillary fiber construct A standard hydrated collagen gel was prepared as described above. Fine glass fibers (phosphate glass with a diameter of 35 microns) were placed in parallel on the surface of the collagen gel. Another collagen gel was formed on the surface of the composite material, and the entire structure was strengthened by the above-described plastic compression.

  The resulting material was observed to consist of a biomaterial in the form of a sandwich containing live cells and fine guide fibers traversed along its entire length.

Bone biomimetic A standard hydrated collagen gel is prepared as described above except that a plurality of porous beads are dispersed between the fibers of the collagen scaffold (FIG. 11).

  Compressing a standard mineralized tissue biomaterial porous bead suspended in a collagen gel provides a material suitable for use as a bone biomimetic.

When the production thickness of the roll-shaped structure was about 30 μm, the PC collagen sheet was difficult to handle and fragile, which led to the development of the helical assembly stage (FIG. 12). This involved rolling and rolling the sheet around the short axis of the sheet, resulting in a three-dimensional rod-like “tissue” construct. The helically assembled construct was convenient for handling, culturing, and mechanical testing. This method was also used to make a three-dimensional complex meso-scale structure by stacking PC collagen layers to make structures in the range of 10-100 μm.

A standardized helical assembly construct was prepared from an initial gel (with or without cells) 13 mm wide x 33 mm long x about 4 mm thick as described herein. The cell density was initially about 5.8 × 10 ⁵ cells / ml but increased to 5.8 × 10 ⁷ cells / ml after compression. After reducing the thickness to 30-50μm by plastic compression, the sheet is wound along the shorter (13mm) edge to form a spirally assembled rod, 13mm long and 1.75mm diameter construction was gotten. The SEM (bar = 200 μm) across the edge of this construct is shown in FIG.

  In addition, histological analysis of the helical assembly construct indicated that the cells were viable, viable and active many weeks after the construct was placed in culture.

Histological analysis of the roll construct The roll spiral construct was formed from a randomly oriented collagen gel (ie, prior to alignment under tension). The structure and cellular properties of these constructs were histologically analyzed at various times using standard techniques (ie normal formalin fixation in waxes, wax embedding, sectioning (parallel and perpendicular to the long axis). Both), and H & E staining).

  The new tissue at zero time (immediately after compression) contained dense, evenly spaced cells with a normal, viable appearance. The compression process also resulted in smaller structures (ie, <10 μm) that did not exceed 30 μm sheets. These “structures” consisted of lamellae of compressed collagen fibrils parallel to the liquid discharge surface (FLS), in this case the basal plane. This resulted in a directional liquid drain mainly from a single surface. In practice, the individual materials (collagen fibrils and cells) were retained at the discharge interface rather than the compression of the material at the filtration surface. A fine collagen lamella (thickness <1 μm) was evident throughout the depth of the compressed collagen sheet, but the liquid discharge surface so that the overwhelmingly highest and highest density layer was observed in the birefringence pattern In parallel, it was formed.

  The cells in the compressed sheet were observed to be aligned with these lamellae. Prior to helical assembly, the underlying cells appeared normal and active and spread throughout the collagen substrate after 3 hours of culture.

After 0 day and 14 day culture (floating in DMEM medium containing 10% fetal calf serum, 37 ° C, 5% CO ₂ incubator), these cells spread and adhere to the matrix and are active in a uniform distribution It exhibited a normal three-dimensional appearance of fibroblasts in the tissue, with no apparent reduction in sex, viability or cell death. At 14 days, the cells were elongated or discoid within the collagen lamella of the roll-like construct.

  After 5 weeks of culture (unrestrained), a dense distribution of viable cells was observed throughout the thickness of the construct (ie, no cell-free core). Signs of new tissue formation by the underlying cells are observed on the surface of the construct and at its edges, with a concomitant stratification and structural area being observed, resulting in the core matrix and cells generally being Although oriented in the long axis direction, the outer cell layer was more circumferential (confirmed by comparison of cross-sections of parallel and vertical surfaces of the same construct). At this stage, the ends of the construct were relatively irregular and disordered protrusions. Clear cellized lamellae (with channels in between) were not evident when they were formed, but at this stage they were evident between obvious tissue layers. Prominent stratification of collagen was evident in the sheet surrounding the construct, with the underlying flattened discoid cells inside the collagen layer. Thus, the 5 week construct was morphologically equivalent to the early natural connective tissue, with healthy cells embedded within and between the layers of collagen fibers.

  Thus, the cells were clearly viable and attached within the dense three-dimensional matrix over the entire culture period. The structure of the matrix gradually changed (in terms of apparent cell / matrix growth across the surface, as well as the formation of channels within the core) during long-term culture, but throughout the entire construct (ie including the core) The high-density connective tissue surrounding active living cells that distributed was generally unchanged.

Mechanical properties of the roll-like constructs The constructs have been demonstrated to be viable and cellized (having a dense matrix) for days and weeks after construction, so that the mechanical properties of the constructs are Were tested as follows.

  FIG. 14 shows a typical tensile test stress-strain curve obtained from a cell-free (zero hour) helical assembly construct, which shows a corrected plot of material behavior and modulus of elasticity to failure. . The lines of the construct indicate the toe region, the linearly extending region, and the yield point (which leads to the failure point). Cellular gels were observed to show a decrease or disappearance of the toe region and a decrease in fracture stress and strain compared to cell-free gels.

  The cellular elements of the cellular gel constituted the majority of the total volume of the construct, which was clearly weaker than the compressed collagen matrix.

  The cell-free construct made by the standard PC method (50 g, 5 minutes) exhibits a functionally useful average tensile fracture strength (0.6 MPa ± 0.11) and an elastic modulus of 1.5 MPa ± 0.36. There was no significant change after 2 weeks in culture. The fibroblast seeded construct resulted in high cell density after PC, which reduced yield strength and elastic modulus by 71% and 34%, respectively, as expected (FIG. 15) and eliminated the toe region at the time of zero hour culture. (FIG. 16). The elastic modulus decreased by 1/3, but the fracture stress hardly changed with the cell seed (0.55 MPa ± 0.06).

  Figures 15 and 16 show a summary of mean breaking strength and modulus (mean +/- SD), but for both cellular and cell-free constructs at zero time and for cellular constructs at 2 weeks.

  After 2 weeks in culture, the material properties of the cellular construct improved to almost that of a cell-free construct, providing clear evidence of functional cell remodeling of the new collagen scaffold. Both yield stress and modulus increased for the cellular construct at time zero (approximately 1.5 and 2 times, respectively), but the fracture stress remained the same. Significant toe strain reappeared (-2%-which was almost lost from the cell-free gel) and yield strain increased from 13 to 23%.

  These data relate to constructs made without tensile pre-alignment of collagen fibrils. The strain to failure would decrease dramatically (from a high level of 50%) for the pre-aligned construct, but in this construct the load was parallel to the direction of the collagen fibers.

Double compression The collagen roll (ie, helical) construct made as described above was subjected to a second compression (secondary compression) using a similar technique, resulting in a flat strap of dense collagen material. .

  Cell and collagen density was found to be proportional to fluid loss. The average weight loss (and thus liquid discharge) from the second compression was approximately 60% (of the helical construct).

  The mechanical properties of the construct were tested as described above and the results are shown in FIG. It was found that the fracture stress approaches 2 MPa, the elastic modulus is almost equal to 2.5 MPa, and the failure strain is 70% at maximum. Therefore, secondary compression has been found to further improve the mechanical properties of the construct.

  Compression in the orthogonal plane of the helical rod (ie parallel to the long axis) resulted in similar compression properties and the material resulted in a flattened disc of collagen.

It has been found that nano-micro-scale structures and heterogeneous layer formation two more levels of organization are controllably generated within the three-dimensional structure by the methods described herein. Lamella formation was predicted for the filtration of liquid from the “liquid discharge surface” and was first observed by birefringence under polarized illumination. Lamella was also observed by TEM and could be measured as the difference in average collagen fibril density between the matrix edge and the center (FIG. 18). The lamella was always parallel to the liquid discharge surface. Not only was there a 7-8 fold increase in average fibril density between the stable control and the PC-treated construct, but there was also a significant increase in fibril density between the edge and center of the construct. .

  Various changes in the orientation and position of the mesostructure (lamellar) that can be created by controlling the liquid discharge surface and the liquid discharge direction during plastic compression are shown in FIG.

  Directional collagen lamella is formed adjacent to and parallel to the liquid discharge surface (in contact with the porous layer), but is rarely formed adjacent to the impermeable layer, Total asymmetry results (FIGS. 19A and B). The lamella can be compressed using a second flow vector orthogonal to the initial direction (FIG. 19C). This can result in the orientation (shown in FIG. 19D) of parallel (uniaxial) fiber bundles from the compressed lamella formed in FIG. 19C. In FIGS. 19A and 19B, the collagen lamella is composed of fibrils packed parallel to the discharge surface, but is considered to be oriented in a random direction on the orthogonal surface. Simultaneous biaxial liquid removal results in two symmetrical planes, as shown in FIG. 19D, and further compresses the collagen lamella perpendicular to the direction shown in FIG. 19B, resulting in a structure equivalent to the natural collagen fibril bundle. Can occur. A combination of a wicking layer and a single plane compressive load can be used to control the flow rate.

  The effect of perturbation localized in the liquid flow vector is shown in FIG. FIG. 20A shows the improvement of the porous blotting layer (in this case the upper surface of the gel), which contains alternating porous [↑] and nonporous [×] strips, but these strips are pressed onto the gel surface (Nonporous patches can be made, for example, by applying varnish). The liquid under the nonporous patch is forced to flow parallel to the discharge surface and tends to form a “partial” lamella that is orthogonal to the surface. However, under the porous portion, the liquid flow deviating from its original location merges with the main, vertically outgoing flow to form a normal lamella parallel to the discharge surface. This creates an orthogonal pattern of lamella (with a transition zone) and can effectively create channels across the entire collagen sheet on the surface. Such channel formation has a guiding ability to pre-pattern growth into nerves, blood vessels or epithelium. Different non-porous patterns (eg, discs) can be used, for example, in combination with a second orthogonal discharge surface to adjust the lamella and channel structure (FIG. 20B). Again, compression is applied to one plane to control the flow rate.

  The helical assembly incorporated this meso-scale structure into the second stage of spatial organization at a scale of several hundred microns. Composite PC constructs were made by adding either fibronectin (Fn) or hydroxyapatite (HA) particles. The hydrated Fn material also received plastic compression, albeit slightly more than collagen. Fibrous Fn layers are compressed together toward the collagen, the two layers are rolled together in an offset spiral, creating an outer collagen tube or sheath around the Fn core, and the layers are well integrated . The structural complexity of forming zones along the length of the construct was controllable, as shown by forming a “hard tissue” end to the construct. This was achieved by placing a layer of HA particles along the long side of the gel surface before compression. By winding this structure around a short side, a “ligament model” with hard nodes on both ends of the collagen cable was generated. A cross-sectional view of the composite (FIG. 21) shows HA particles packed between dense collagen layers. Hydroxyapatite (HA) end zones are not only capable of anchoring to bone (eg, screwed) in vivo, but can also bind very well to mechanical bioreactors.

  Such an example with or without cells reveals the potential for spatial and compositional ranges for fabricating meso-scale structures.

Complex Assembly As noted above, multiple layers of prefabricated thin cellular sheets with a range of properties and contents can be assembled. Such elements are themselves prefabricated, in part, using plastic compression to dominate in scale strength and texture composition. These elements take advantage of the helical assembly approach to achieve micro (meso) -scale spatial biological complexity, with final compression joining them together. For example, for nerve repair implants, a sturdy outer sheath with a non-adhesive outer surface, built-in on and off ramps (cell (neurite) survival, ingrowth and subsequent ingrowth, and reintegration) A graded growth factor content microdepot is incorporated into the core portion.

  FIG. 23 shows a concentric helical assembly of such a neural implant. The concentric axis defines the regeneration growth axis (arrows in both directions). Each element of the final construct, or functional mode (sheath, axon guide filament, sustained release depot, special lamp, etc.) is made as a sheet (thickness <50 μm, ie cell size) or as an intersheet fiber insert The An example of a cut-out view visible inside encapsulates the outer cell seed layer [1] (ideally with a sustained release of hyaluronan to prevent attachment), followed by growth factors etc. (depots with different compositions) A growth factor depot impregnated sheet [2]. There is also a sheet of traditional fibrous axon guide “deck” material (currently fibronectin) [3]. The end layer [2] may have an appendage [4], which forms a device off-ramp (eg, a tactile mesomaterial).

  When this composite material is rolled from the bottom (arrows and insets) as shown in the figure, multiple meso-scale elements automatically form the required spatial arrangement (partial overlap, order, boundary Surface, having a gap). It also has the great advantage (using thin layers) of diverting two-dimensional nanolayer formation technology for use in macro devices with complex three-dimensional spatial / directional structures. This development requires the plotting and monitoring of spatial patterns formed at the nano- (electron microscope) level.

In vivo transplanted helical constructs were prepared by plastic compression with or without embedded rabbit tendon fibroblasts (allogeneic) as described above. The PC implant group was sutured between the rabbit intercostal muscle layers and secured tightly between the intercostals so that it was pulled as the animal breathed. If the implant is removed after 1, 3 or 5 weeks and the implant is still in place and intact, the amount and pattern of cell proliferation, collagen deposition (birefringence), identification of inflammatory cells, capillary growth Was analyzed histologically. Capillary growth was also monitored by backscatter spectroscopy to detect changes in absorbance associated with (oxy) hemoglobin and to monitor hemoglobin and oxygenated hemoglobin levels at the time of sacrifice. Finally, the collected constructs were examined for mechanical strength in vivo over time (using the DMA tensile test apparatus as described above).

  Cells grown for a week were moderate in cell-free constructs, but more in cellular constructs, growing to deeper materials, mainly on the surface and at short distances around the helix. There were few signs of new collagen deposition and no blood vessel growth.

  The 3 week cellular construct was macroscopically more vascularized and hardly grew into the cell free construct. Cell ingrowth was still greater for cellular constructs, and the deposition of collagen in the multicellular layer was significant, covering the outer surface and the deeper intersheet helical surface. The direct involvement of the seeded cells appeared to be very small because they are still deep in the PC collagen material, but the total activity is histologically less than that of the cell-free construct. The seeded construct was consistently larger.

  At 5 weeks, there was a further increase in growing cells for both types, completely occupying the core of the construct on which the cells were seeded. The cell-free construct produced good angiogenesis by this stage, and in each type, a lot of collagen was newly deposited between the helix and over the surface.

  Backscattering spectroscopy showed a rapid increase in the hemoglobin index between 1 and 3 weeks, and was 3 times greater at 3 weeks for cellular and cell-free constructs (Figure 24). The cell-free construct caught up in the next 2 weeks. This indicates that the cell construct was much more vascularized at 3 weeks than the construct without seeded cells.

  Cell infiltration was quantified and increased significantly at each time point, but the difference between cell type and cell-free type did not reach statistical significance. Cd31 staining increased after 3 weeks in all constructs (FIG. 25). CD 45% leukocyte staining (FIG. 26) was relatively constant throughout. Total cell density increased slightly between 1 and 3 weeks (significantly increased only for cell constructs), but increased rapidly between 3 and 5 weeks (overall> 50%) (overall) Figure 27).

  The mechanical strength of the construct was measured after recovery. The failure stress of the cell-free construct first decreased during the first 3 weeks and then began to recover. In contrast, cellular constructs increased in vivo for 3 to 5 weeks without weakening (Figure 28). The total strain to failure was greater for the cellular construct and increased slightly (NS) during the implant period (Figure 29). The elastic modulus was greater in the cell-free than the cellular pre-implant construct. The elastic modulus decreased from 50 to 30% of the pre-implant value in all implant constructs. That is, the implanted construct became more rigid (Figure 30).

  The cellular construct increased significantly in diameter from 0 to 5 weeks, but not in the cell-free construct, and the cellular construct at 5 weeks was significantly larger than the cell-free construct at the same time. Clearly, the initial loss of mass was at least reversed in any construct (FIG. 31).

Summary The method described herein is a radically new method for rapidly creating new tissues and bioartificial materials with controllable meso-scale structures. The simplicity and extreme rapidity of this method are its key features. This method is expected to immediately impact a variety of biomedical fields from tissue engineering, regenerative and reconstructive surgery to sustained drug delivery materials.

Fabrication by plastic compression (PC) results in a reproducible and controllable structure with tensile properties approximately equal to those of natural tissue. The structure is still often weaker than the strongest tissue (eg, about 100 MPa ¹⁵ for tendons), but can now be sutured and loaded in vivo for further remodeling. Further enhancement of the mechanical properties is brought about by subjecting the rolled or folded structure to a secondary compression, for example parallel to the long axis of the rolled structure. Such controlled secondary compression has been shown to reduce the liquid content by a further 60% but increase the fracture stress reproducibly up to 100%.

  Control of the mechanical properties of the construct is possible by selecting the compression method used and the extent (or stage) at which the interstitial fluid is drained. Stage 1 is believed to produce a mechanically fragile material due to slow liquid loss during self-compression. Stage 2 is represented by using one active compression and spiraling (50 g, 5 minutes), while stage 3 can be achieved by a second compression applied to the helical construct itself. In fact, rapid primary compression can be delayed by, for example, compressing with only the wicking layer in order to better control the removal of the liquid and thus the thickness of the collagen sheet.

  Meso-scale fabrication of structural features is accomplished by collagen lamella formation, surface topography, zoning, and layering by helical assembly, but these fabrications engineer local biomimetic structures. It addresses the pressing need to create. For example, it can be used to provide clearly differentiated ingrowth, intermediate growth and ingrowth structures for peripheral nerve regeneration (HM Geller et al. Exp. Neurol. 2002). , 174,125). Two examples of helical assemblies shown here are neural regeneration templates (collagen sheath and Fn neurite guidance core (J. Phillips et al., Biomaterials 2003, 25, 2769)) and ligament models (collagen cable and anchorage For hard tissue edges).

  In vitro constructs clearly undergo active remodeling even in the simplest culture systems, and mechanical loading further improves tensile properties. The cells exhibited a tissue-like distribution and morphology and were aligned along a number of collagen lamellas in a flat spindle or disc shape. Such channeling between the collagen layers aids perfusion and improves the tissue-like structure.

  None of the constructs subjected to mechanical testing were pre-tensioned and aligned, so collagen is essentially non-directional in the XY plane (ie, the fibril orientation in the lamella is in the 'Z' plane) ), Which is thought to have contributed to relatively high compliance. Induction of XY fibril anisotropy results in a material that is more similar to natural tissue, particularly with respect to stiffness relative to the fibril alignment axis. In addition, the fracture and yield stresses in this study used cross-sectional areas measured with large specimens. This means that the fracture strength has been underestimated, as it is clear that the hand spiraling process leaves some channeling between the layers. The finding that the mechanical properties were improved during culture demonstrates that the collagen matrix has undergone natural cell remodeling, suggesting that the construct is functionally integrated in vivo.

  The above experimental data show that the PC fabrication described herein: (a) does not substantially reduce the viability of the underlying cells; (b) a collagen matrix with effective mechanical properties (strength and compliance) (C) (i) a composition having a full range of cell densities and matrix densities, and (ii) spatial / zone heterogeneity and structural differences adapted to the exact biomimetic structure. It can be controlled as expected so as to produce a directionality. The finding (a) that relatively high fluid shear (ie, <100 minutes contraction in <5 minutes) is tolerated by the underlying fibroblasts is the current use in tissue fabrication using this process for ultrafast fabrication. Indicates that obstacles can be avoided. Preparation of a matrix having mechanical strength such as tissue (b) means that the construct can be transplanted directly as it is or after a short incubation period. Furthermore, the ability to limit the degree of compression with good reproducibility means that the size and mechanical properties can be controlled. So far, with weak tensile properties, collagen gel stitching has been difficult or impossible without chemical cross-linking or reinforcement with other materials, which has reduced the mimicry of natural matrices.

  The advantages of not requiring the involvement of the underlying cells are: (i) the possibility of PC production at low temperatures (minimizing hypoxia and nutrient requirements); (ii) promoting differentiation of stem or progenitor cells And (iii) the production of cell-free biomaterials with complex meso-structures (P. Vadgama et al., In Institute of Materials, Mineral and Mining, London (report number: FMP / 03/01 / IOM3), 2004, pp. 14-18). Other types of hydrated gels such as fibronectin, fibrin, or synthetic polymers can be used to provide other meso-scale organization. Furthermore, such a process is suitable for large scale production or bedside “individual” production.

参考文献
M.Eastwoodら Biophys. Biochim. Acta. 1994:1201;186-192.
Eastwood M.ら J. Cell Physiol. 1996:166;33-42.
Cacou C.ら Cell Eng. 1996:1:109-114.
Brown RAら J. Cell Physiol 1996:169;439-447.
Talas G.ら Int. J. Biochem. Cell Biol.1997:29;261-70.
Brown RA.ら J. Cell Physol. 1998:175;323-32.
M.Eastwoodら Cell Motility Cytoskel. 1998:49;13-21.
Eastwood MEら J. Institute Mech. Eng. 1998. 212;85-92.
RA. Porterら Wound Repair Regen. 1998:6;157-166.
Mudera VC. らCell Motility Cytoskel. 2000:45;1-9.
Prajapati RTら Wound Repair Regen. 2000:8;227-238.
Parjapati.RTら Wound Repair Regeneration. 2000:8;239-247.
Brown RAら (2002) Exp. Cell Res. 274:310-322..
Sethi, KKら (2002). Cell Motil. Cytoskel. 52: 231-241.
Cheema U.ら 2003. Cell Motil. Cytoskel. 54:226-236.
Sethi KKら 2002. Wound Rep. Regen. 10; 397-408.
Brown RA. Next Generation Tissue Engineering: Clinical Applications and Mechanical Control. In ‘Future Strategies for Tissue and Organ Replacement’ (Polak JM., Hench LL., Kemp P.編集). World Scientific Publishing, Singapore. (2002) 48-75.
Tomasek JJ.ら (2002) Nature Reviews Molec. Cell Biol. 3:349-363.
Brown RA.ら Wound Rep. Regen. 1997:5;212-21.
Krishnan L., Weiss J A, Wesserman M D, Hoying J B. (2004) Tissue Eng. 10:241-252
D. Huangら, Ann. Biomed. Eng. 1993, 21, 289.
F. Volrath, D. P. Knight, Nature 2001, 410, 541.
Z. Fengら, Artificial Organs 2003, 27, 84
I.V. Yannasら Science 1982, 215, 174.
S.F. Badylak, Rら J. Biomed. Mater. Res. 1995, 29, 977.
A. Curtisら, Arch. Dermatol. 1998, 134, 293.
D. C. Wallace, J. Rosenblatt, Adv. Drug Delivery Rev. 2003, 55, 1631
Z. Rusczcak, W. Friess, Adv. Drug Delivery Rev. 2003, 55, 1679.
J. Garvinら Tissue Eng. 2003, 9, 967.
S. Calveら Tissue. Eng. 2004, 10, 755.
H. Schechtman, D.L. Bader. J. Biomech. 1997, 30, 829.
H.M. Geller, J.W. Fawcett, Exp. Neurol. 2002, 174, 125.
J. Phillips, R.A. Porter, Z. Ward, S. Standring, V. King, R.A. Brown, Biomaterials 2003, 25, 2769.
P. Vadgamaら in Institute of Materials, Mineral and Mining, London (report number: FMP/03/01/IOM3), 2004, 14-18.
A.M. Burt, D.A. McGrouther, in Animal cell biotechnology, Academic Press, New York 1994 p 150.
R.A. Brownら Biomaterials 1994, 15, 457.
L.C. Junqueiraら Histochem J. 1979, 11, 447.
A.R. Spurr, J. Ultrastruct. Res. 1969, 26, 31.
L. Wollweberら J. Microsc. 1981, 121, 185

References
M. Eastwood et al. Biophys. Biochim. Acta. 1994: 1201; 186-192.
Eastwood M. et al. J. Cell Physiol. 1996: 166; 33-42.
Cacou C. et al. Cell Eng. 1996: 1: 109-114.
Brown RA et al. J. Cell Physiol 1996: 169; 439-447.
Talas G. et al. Int. J. Biochem. Cell Biol. 1997: 29; 261-70.
Brown RA. Et al. J. Cell Physol. 1998: 175; 323-32.
M. Eastwood et al. Cell Motility Cytoskel. 1998: 49; 13-21.
Eastwood ME et al. J. Institute Mech. Eng. 1998. 212; 85-92.
RA. Porter et al. Wound Repair Regen. 1998: 6; 157-166.
Mudera VC. Et al. Cell Motility Cytoskel. 2000: 45; 1-9.
Prajapati RT et al. Wound Repair Regen. 2000: 8; 227-238.
Parjapati.RT et al. Wound Repair Regeneration. 2000: 8; 239-247.
Brown RA et al. (2002) Exp. Cell Res. 274: 310-322 ..
Sethi, KK et al. (2002). Cell Motil. Cytoskel. 52: 231-241.
Cheema U. et al. 2003. Cell Motil. Cytoskel. 54: 226-236.
Sethi KK et al. 2002. Wound Rep. Regen. 10; 397-408.
Brown RA. Next Generation Tissue Engineering: Clinical Applications and Mechanical Control. In 'Future Strategies for Tissue and Organ Replacement' (Edited by Polak JM., Hench LL., Kemp P.). World Scientific Publishing, Singapore. (2002) 48- 75.
Tomasek JJ. Et al. (2002) Nature Reviews Molec. Cell Biol. 3: 349-363.
Brown RA. Et al. Wound Rep. Regen. 1997: 5; 212-21.
Krishnan L., Weiss JA, Wesserman MD, Hoying J B. (2004) Tissue Eng. 10: 241-252
D. Huang et al., Ann. Biomed. Eng. 1993, 21, 289.
F. Volrath, DP Knight, Nature 2001, 410, 541.
Z. Feng et al., Artificial Organs 2003, 27, 84
IV Yannas et al. Science 1982, 215, 174.
SF Badylak, R et al. J. Biomed. Mater. Res. 1995, 29, 977.
A. Curtis et al., Arch. Dermatol. 1998, 134, 293.
DC Wallace, J. Rosenblatt, Adv. Drug Delivery Rev. 2003, 55, 1631
Z. Rusczcak, W. Friess, Adv. Drug Delivery Rev. 2003, 55, 1679.
J. Garvin et al. Tissue Eng. 2003, 9, 967.
S. Calve et al. Tissue. Eng. 2004, 10, 755.
H. Schechtman, DL Bader. J. Biomech. 1997, 30, 829.
HM Geller, JW Fawcett, Exp. Neurol. 2002, 174, 125.
J. Phillips, RA Porter, Z. Ward, S. Standring, V. King, RA Brown, Biomaterials 2003, 25, 2769.
P. Vadgama et al. In Institute of Materials, Mineral and Mining, London (report number: FMP / 03/01 / IOM3), 2004, 14-18.
AM Burt, DA McGrouther, in Animal cell biotechnology, Academic Press, New York 1994 p 150.
RA Brown et al. Biomaterials 1994, 15, 457.
LC Junqueira et al. Histochem J. 1979, 11, 447.
AR Spurr, J. Ultrastruct. Res. 1969, 26, 31.
L. Wollweber et al. J. Microsc. 1981, 121, 185

The plastic compression by mechanical compression of (a) isotropic gel and (b) anisotropic gel is shown. Fig. 2 shows a combination of plastic compression and blotting of a gel. For horizontal and 45 ° tilted gels, shows a graph of gel mass reduction (ie% liquid reduction) with respect to gel self-compression due to the weight load of the gel itself with increasing gel time before compression (dots mean ± SD, n = 3). Figure 5 shows moisture loss from constructs solidified in air. Scanning electron micrographs of compressed constructs without tension pre-alignment with 25% uniaxial strain (upper figure) and after tension pre-alignment (lower figure) are shown (white bars are 100 μm in both pictures). The relative importance of liquid uptake (absorption) by paper and nylon porous layers is shown. Figure 6 shows the effect of compression of a homogenous collagen gel under a range of dead load and compression time conditions. In contrast to PC-free and control cell-free constructs, the difference in collagen fiber density between the edge and center of the specimen sheet is evaluated by image analysis of transmission electron micrographs (note: control) The gel was self-compressed to 57%; the PC construct was 93%). Box-whisker plots show median fibril density in the full range (bars) and interquartile range (boxes) for the edge and center of the control and PC constructs [average control value is approximately 1% (The difference is not significant) and PC constructs were 9.9% (edge) and 6.6%, respectively (p <0.005) n = 13-17]. The fibril density increased 7-8 fold in each region between the PC and control constructs (* and ** p <0.0001). Fig. 4 shows an arrangement for generating a compression impulse from the core of the structure to provide plastic compression. Figure 2 shows a conical flow chamber for continuous flow extrusion-plastic compression. FIG. 5 shows plastic compression of a suspension of standard mineralized tissue biomaterial porous beads added into a collagen construct as a biomimetic form of bone. Figure 2 shows an overview of the preparation of a PC construction sheet and the helical assembly of the PC construction sheet into a three-dimensional structure, showing the lamellar structure of the edge and the spiral lamella after winding on the three-dimensional structure. Scanning electron micrograph of an end view of a helical structure, showing a collagen sheet layer (with interface shrinkage: about 2 mm in diameter). Figure 2 shows a typical tensile test stress-strain curve obtained from a cell-free (zero hour) helical assembly construct, which shows the behavior of the material to failure and a corrected plot of modulus. The lines of the construct indicate the toe region, the linearly extending region, and the yield point (which leads to the failure point). Shows bar graphs (mean ± SD; n = 6-8) of toe, yield point, maximum breaking stress and elastic modulus for cellular and cell-free helical constructs. Shows bar graphs for strain values at toe region, yield point, and maximum failure point. Test constructs were seeded with human skin fibroblasts (HDF) cultured for zero hours or two weeks, or were cell free. 2 shows the tensile stress-strain behavior of collagen gel subjected to secondary compression. A SEM view of the multilamellar structure inside a single (30 μm) PC collagen sheet is shown, the thickness is typically <1-2 μm. Fig. 4 shows a diagram representing variations of mesostructures (lamellas) that can be created by controlling the liquid discharge surface and the liquid discharge direction during plastic compression. The open plate represents a porous absorbent surface, and the shaded plate represents an impervious sheet that does not allow liquid flow. Figure 5 shows the effect of localized perturbations in the liquid flow vector. The open plate represents a porous absorbent surface, and the shaded plate represents an impervious sheet that does not allow liquid flow. It is a scanning electron micrograph of an end view of a helical structure showing a collagen sheet layer reinforced with hydroxyapatite particles (with interface shrinkage: about 2 mm in diameter). Scanning electron micrograph of an end view of a helical construction showing a collagen sheet layer reinforced with soluble glass fibers (arrows) (with interface shrinkage: about 2 mm in diameter). FIG. 4 shows a schematic diagram of a general “stratified” design for concentric helical assemblies of complex neural implant constructs. Figure 2 shows the hemoglobin index for the duration of implantation for in vivo implants. The percentage of total cells that were CD31 positive in the in vivo implant is shown. The percentage of total cells that were CD45 positive in the in vivo implant is shown. Shown is the average number of cells per slide in an in vivo implant. The change in fracture strain with transplantation from the 0 hour construct is shown. Cell-free constructs are shown as dark diamonds and cellular constructs are shown as light squares. The change in destructive force with transplantation from the 0 hour construct is shown. Cell-free constructs are shown as dark diamonds and cellular constructs are shown as light squares. The average elastic modulus of an in vivo implant is shown. The diameter of the construct when recovered after transplantation is shown (dark color is acellular, light color is cellular).

Claims (46)

Providing a gel comprising a scaffold fiber matrix and interstitial fluid , further comprising viable cells in the interstitial fluid; and
Apply external plastic compressive force to the gel to drain 50% to 99.9% of the interstitial fluid from the gel in less than 10 minutes,
A method for producing a biomaterial, comprising plastically compressing the gel thereby producing a biomaterial containing viable cells.   Surviving cells are muscle cells, hepatocytes, kidney cells, heart cells, lung cells, gastrointestinal cells, bronchial cells, eye cells, germ cells, vascular cells, nerve cells, secretory cells, stem cells, fibroblasts, Schwann cells, 2. The method of claim 1, wherein the method is selected from the group consisting of smooth muscle cells, endothelial cells, urothelial cells, bone cells, chondrocytes, and tendon cells.   The method of claim 1 or 2, wherein the cells are cultured in the gel for less than 6 hours prior to compression. 4. The method according to any one of claims 1 to 3, wherein the gel has a cell density of at least 1 x 10 < ⁵ > / ml. The method of claim 4, wherein the compressed biomaterial has a cell density of at least 1 × 10 ⁷ cells / ml. The gel is compressed in aqueous liquids, the method according to any one of claims 1-5. Interstitial fluid is discharged through the liquid discharge surface of the gel, the biomaterial comprises one or more lamella aligned with the liquid discharge surface, The method according to any one of claims 1 to 6. 8. A method according to any one of the preceding claims, wherein interstitial liquid is removed from the gel by applying mechanical force to the gel. The method of claim 8 , wherein the mechanical force is a compressive force. For aligning the scaffold fibers and / or cells further comprising adding a tension to the gel, the method according to any one of claims 1-9. 10-50% strain is added to the gel, the method of claim 10. 12. A method according to claim 10 or 11 , wherein tension is applied prior to plastic compression. The method according to any one of claims 1 to 12 , wherein the gel comprises a capillary filament. 14. The method of claim 13 , wherein capillary filaments are soluble and lysis of the filaments results in capillary channels in the biomaterial. The method according to any one of claims 1 to 12 , wherein the gel further comprises porous beads, and the plastic compression pushes the scaffold fibers into the pores of the porous beads. 16. A method according to any one of the preceding claims, wherein the plastic compression forms a sheet of biomaterial. 17. A method according to claim 16 , comprising rolling the sheet into a roll to form a roll of biomaterial. 17. The method of claim 16 , comprising folding the sheet to form a folded biomaterial having two or more layers. 19. A method according to claim 17 or 18 , comprising plastically compressing the rolled or folded biomaterial. The method according to any one of claims 1 to 19 , wherein the scaffold fiber is a collagen fiber. 21. The method according to any one of claims 1 to 20 , wherein the interstitial fluid is a cell culture medium. 22. The method of any one of claims 1-21, comprising storing the biomaterial under conditions that maintain the viability of the cells but do not support cell growth.   23. The method of claim 22, comprising storing the biomaterial at 0-5 ° C. A biomaterial produced by the method according to any one of claims 1 to 23 . 25. A tissue equivalent implant comprising or consisting of the biomaterial of claim 24 . A tissue equivalent implant comprising a plastically compressed gel, the gel comprising a scaffold fiber matrix and interstitial fluid, the interstitial fluid comprising viable cells , and compression
Providing a gel comprising a scaffold fiber matrix and interstitial fluid, further comprising viable cells in the interstitial fluid; and
Applying an external plastic compressive force to the gel that drains 50% to 99.9% of the interstitial fluid from the gel in less than 10 minutes
The tissue equivalent implant obtained by the method comprising . 27. The tissue equivalent implant of claim 26 , wherein the plastically compressed gel is a sheet wound in a roll. 27. The tissue equivalent implant of claim 26 , wherein the plastically compressed gel is a folded sheet. 29. A tissue equivalent implant according to any one of claims 26 to 28 , wherein the folded or rolled sheet is further subjected to plastic compression. Placing overlapping the second gel on the gel, however, the second gel may comprise a Matrigel box and interstitial fluid of the scaffold fibers, and
Applying an external plastic compressive force that drains 50% to 99.9% of the interstitial fluid from the gel in less than 10 minutes to produce the biomaterial, where the gel forms separate layers in the biomaterial thing,
The living body of claim 1, wherein the one or more gels comprise viable cells, and the one or more gels are plastically compressed to produce a biomaterial comprising viable cells. Material production method. 32. The method of claim 30 , wherein the gel is a different region of the same gel. 32. The method of claim 30 , wherein the gel is a different gel. 35. The method of claim 32 , wherein two or more gels are placed on top of each other. 34. The method of claim 33 , wherein the two or more gels differ in one or more of the following features: fiber organization in the scaffold material, density of the scaffold in the gel, fiber type. 35. A method according to any one of claims 30 to 34 , wherein the plastic compression forms a sheet of biomaterial having two or more layers. 36. The method of claim 35 , comprising folding the sheet to form a folded biomaterial. 36. The method of claim 35 , comprising winding the sheet into a roll to form a roll of biomaterial. The roll has the following characteristics:
The space between the layers,
Partial overlap between adjacent layers, and different combinations or permutations of layers at different axial positions along the roll,
38. The method of claim 37 , wherein the method is a helical assembly having one or more of: 39. A method according to any one of claims 36 to 38 , comprising plastically compressing a rolled or folded biomaterial. 40. A method according to any one of claims 30 to 39 , comprising shaping and / or shaping the biomaterial. 41. A method according to any one of claims 30 to 40 , wherein the scaffold fiber is a collagen fiber. 42. The method according to any one of claims 30 to 41 , wherein the interstitial fluid is a cell culture medium. Surviving cells are muscle cells, hepatocytes, kidney cells, heart cells, lung cells, gastrointestinal cells, bronchial cells, eye cells, germ cells, vascular cells, nerve cells, secretory cells, stem cells, fibroblasts, Schwann cells, 43. The method according to any one of claims 30 to 42 , selected from the group consisting of smooth muscle cells, endothelial cells, urothelial cells, bone cells, chondrocytes, and tendon cells. A biomaterial produced by the method according to any one of claims 30 to 43 . 45. A tissue equivalent implant comprising or consisting of the biomaterial of claim 44 . 46. Use of a tissue equivalent implant according to any one of claims 25 to 29 and 45 in the manufacture of a medicament for use in the treatment of damaged tissue.

Images